Diagnostic imaging techniques, such as magnetic resonance imaging (MRI), x-ray, nuclear radiopharmaceutical imaging, optical (ultraviolet, visible and/or infrared light) imaging, and ultrasound imaging, have been used in medical diagnosis for a number of years. In some cases, the use of contrast media to improve the image quality or to provide specific information has been ongoing for many years. In other cases, such as optical or ultrasound imaging, the introduction of contrast agents is imminent or recent.
MRI and optical imaging methods are unique among imaging modalities in that they yield complex signals that are sensitive to the chemical environment and state of the targeted tissue. While the signal from x-ray or radionuclide agents remains the same whether the agents are free in plasma, bound to proteins, or trapped inside bone, certain agents for MRI and optical imaging will have different signal characteristics in differing physiological environments and pathological states. For example, by binding to tissue components, MRI contrast agents can show changes in the induced relaxation rates or chemical shifts of nearby or attached nuclei. Similarly, an optical dye may exhibit changes in its absorbance, reflectance, fluorescence, phosphorescence, chemiluminescence, scattering, or other spectral properties upon binding.
In general, to provide diagnostic data, the contrast agent must interfere with the wavelength of electromagnetic radiation used in the imaging technique, alter the physical properties of tissue to yield an altered signal, or, as in the case of radiopharmaceuticals, provide the source of radiation itself. Commonly used materials include organic molecules, metal ions, salts or chelates, including metal chelates, particles (particularly iron particles), or labeled peptides, antibodies, proteins, polymers, or liposomes.
After administration, some agents non-specifically diffuse throughout body compartments prior to being metabolized and/or excreted; these agents are generally known as non-specific agents. Alternatively, other agents have a specific affinity for a particular body compartment, cell, cellular component, organ, or tissue; these agents can be referred to as targeted agents.
One application for diagnostic imaging techniques has been in the monitoring of interventional therapies. Common interventional therapies include targeting an undesired tissue or tissue component with high thermal energy using focused ultrasound (e.g., Cline et al., “MR Temperature Mapping of Focused Ultrasound Surgery,” Mag. Resn. Med., 31:628–636 (1994)), radiofrequency generators (e.g., Rossi et al., “Percutaneous RF Interstitial Thermal Ablation in the Treatment of Hepatic Cancer,” AJR, 167:759–768 (1996)), microwave antennae (e.g., Schwarzmaier et al., “Magnetic Resonance Imaging of Microwave Induced Tissue Heating,” Mag. Resn. Med., 33:729–731 (1995)), and lasers (e.g., Vogl et al., “Recurrent Nasopharyngeal Tumors: Preliminary Clinical Results with Interventional MR Imaging-Controlled Laser-Induced Thermotherapy,” Radiology, 196:725–733 (1995)); the use of cryoablation (i.e., liquid nitrogen) and the injection of denaturing liquids (e.g., ethanol, hot saline) directly into the undesired tissue (e.g., Nagel et al., “Contrast-Enhanced MR Imaging of Hepatic Lessions Treated with Percutaneous Ethanol Ablation Therapy,” Radiology, 189:265–270 (1993) and Honda et al., “Percutaneous Hot Saline Injection Therapy for Hepatic Tumors: An Alternative to Percutaneous Ethanol Injection Therapy,” Radiology, 190:53–57 (1994)); the injection of chemotherapeutic and/or chaotropic agents into the tissue (e.g., Pauser et al., “Evaluation of Efficient Chemoembolization Mixtures by Magnetic Resonance Imaging of Therapy Monitoring: An Experimental Study on the VX2 Tumor in the Rabbit Liver,” Cancer Res., 56:1863–67 (1996)); and photodynamic therapies, wherein a cytotoxic agent is activated in vivo by irradiation with light (e.g., Dodd et al., “MRI Monitoring of the Effects of Photodynamic Therapy on Prostate Tumors,” Proc. Soc'v Mag. Resn., 3:1368, ISSN 1065–9889 (Aug. 19–25, 1995)). The shared goal of all such interventional therapies is the treatment of undesirable tissue or tissue component (i.e., cancerous, tumorous, neoplastic tissue or tissue component) by causing the necrosis, ablation, coagulation, or denaturation of such tissue.
To obtain the maximum benefit from such interventional methods, and to minimize side effects (e.g., damage to adjacent tissues), it is essential to monitor, in vivo, the efficacy of the therapy. Indeed, to be truly effective, the interventional therapy must continue until the absolute “death” of the undesired tissue or tissue component (nonviability after removal or conclusion of therapy). Thus, one must not only be able to accurately monitor the progress of the therapy, so as to avoid excessive treatment and possible damage to adjacent tissue, but must also be able to accurately distinguish between truly necrotic tissue and those which may have been injured to a certain extent but remain viable nonetheless.
One way to monitor the efficacy of the interventional therapy is to image the undesired tissue or tissue component during or after such therapy. However, any such diagnostic imaging method must be capable of increasing the contrast between tissues of different pathological states (native vs. denatured, viable vs. necrotic) in such a way to provide two basic classes of information:
1) Detection Data. This includes spectroscopic information necessary to determine the pathologic state of the imaged tissue. The ability to provide this class of information relates to the “specificity” and “sensitivity” of the agent.
2) Feedback and Resolution. These classes of information provide the monitoring of interventional therapeutic procedures that destroy or degrade tissue or tissue components. It is envisioned that with some interventional methods, “real time” feedback (about 1–10 seconds) of the therapy's progress is preferred, while with other methods, a post-therapeutic assessment is adequate. With all interventional therapies, precise spatial resolution (about 1–5 mm) of the tissue treated and any effects on surrounding tissues during treatment is desirable.
Current MRI-based methods for monitoring the efficacy of interventional therapies are generally one of two classes: (1) those that do not use an exogeneous contrast agent but rely on some other observable MR parameter (vide infra); and (2) those that use non-specific, extracellular contrast agents. These methods, however, provide virtually no direct information regarding the pathological state of the tissue or tissue component undergoing interventional therapy (e.g., whether it is native or denatured, necrotic or viable). Further, such methods are largely limited to monitoring thermal ablation therapies and provide limited sensitivity to thermally-induced tissue temperature changes.
Several of these MRI-based methods for monitoring thermal ablation therapies rely on temperature-dependent NMR parameters such as relaxation times (T1 and/or T2), the proton resonance frequency (PRF) of water, phase shifts, and the diffusion coefficient. However, these methods suffer from a number of limitations.
For example, one such method involves monitoring the effect of temperature on the T1 relaxation time of tissue. See, e.g., Cline et al., “MR Temperature Mapping of Focused Ultrasound Surgery,” Mag. Resn. Med., 31:628–636 (1994). This approach, however, is inadequate because each tissue has a unique T1 versus temperature profile, and thus, this method requires T1 calibration for each tissue type. The T1 method is also limited in sensitivity, with a tissue dependent change in T1 of only 0.01% to 1.5% per 1° C.
Another method using temperature measurement involves monitoring the effect of temperature on the proton resonance frequency (or chemical shift) of water. This method detects changes in hydrogen bonding and molecular motion of water molecules induced by temperature changes. See, e.g., J. D. Poorter, et al., “Noninvasive MRI Thermometry with the Proton Resonance Frequency (PRF) Method: In Vivo Results in Human Muscle,” Mag. Resn. Med., 33:74–81 (1995). However, the low sensitivity of this method (0.01 ppm/° C.) requires the use of high magnetic field strengths (i.e., >4.7 T) which is clinically undesirable. Further, the determination of the chemical shift of water requires absolute stability of the magnetic field and is also highly dependent upon the magnetic susceptibility of the tissue which varies dramatically among different tissue types. Thus, this method, like the T1 method, also requires extensive calibration for each tissue type. Finally, this method does not provide information regarding thermally-induced tissue necrosis or degradation.
Another known method requires monitoring the effect of temperature on the water proton diffusion coefficient. See, e.g., H. Saint Jalmes, “Precision in Temperature Measurement via T1 or Diffusion Imaging,” Proc. Soc'y Mag. Resn., 2:1072, ISSN 1065–9889 (Aug. 19–25, 1995). This method, however, is also limited because the diffusion coefficient is sensitive to tissue motion and perfusion.
In all of the above methods, physiologic tissue changes due to increased blood flow, tissue metabolism, or induced edema, can result in unpredictable signal variations (i.e., magnetic susceptibility changes). These effects render standard thermal calibration curves to be of little or no value for the accurate monitoring of thermal ablation therapy. Moreover, measuring temperature alone may be insufficient to accurately determine the efficiency of tissue ablation or side effects on surrounding tissues.
Other methods have also been reported which monitor the effect of temperature on the chemical shift of other magnetic nuclei. For example, the cobalt NMR chemical shift is a very sensitive probe of temperature. However, the low receptivity of 59Co requires high field strengths (≧4.7 T), high concentrations, and extensive measuring times. See A. G. Webb et al., “Measurement of Microwave Induced Heating of Breast Tumors in Animal Models Using Cobalt Based NMR,” Proc. Soc'v Mag. Resn., 1:72, ISSN 1065-09889 (Aug. 19–25, 1995). In addition, the toxicity of cobalt agents remains a serious limitation for use in vivo.
Fluorine (19F) NMR has also been used to monitor the temperature-dependent phase transitions of liposome-encapsulated fluorocarbons and fluorinated polymers. See, e.g., Webb et al., “Microencapsulation of Fluorine-Containing Phase Transition Agents for Monitoring Temperature Changes in vivo,” Proc. Soc'y Mag. Resn., 3:1574, ISSN 1065-9889 (Aug. 6–12, 1994). Clinically, however, 19F methods are not useful because of the limited biodistribution of polymeric fluoronated compounds, the chemical shift dependence of fluorinated agents on pH and tissue type, and the need for large magnetic fields. These agents also do not report on thermally-induced tissue necrosis.
Certain contrast agents containing paramagnetic metal complexes have also been suggested to monitor the efficacy of interventional therapies. Such agents can induce large changes in proton chemical shifts (20–40 ppm) of the chelating ligand from the normal range of the water resonance frequency. By paramagnetic shifting of resonances away from the bulk water resonance in vivo, these resonances can be observed. See, e.g., Aime et al., “Yb(III)DOTMA as Contrast Agent in CSI and Temperature Probe in MRS,” Proc. Soc'v Mag. Resn., 2:1109, ISSN 1065-9889 (Aug. 19–25, 1995). Although these hyperfine shifted resonances are temperature dependent, they require the use of high concentrations of the paramagnetic complex and clinically impractical, high magnetic fields to detect temperature changes. These complexes also cannot report on thermally-induced tissue necrosis.
More recently, a method for distinguishing between normal and necrotic liver tissue has been described. Dupas et al., “Delineation of Liver Necrosis Using Double Contrast-Enhanced MRI,” J. MRI, vol. 7, no. 3, pp. 472–77 (1997). This method, however, involves the use of non-specific contrast agents which limits its ability to specifically monitor the state change of the undesired tissue or tissue component. Also, this method requires the administration of multiple contrast agents.
Thus, the known diagnostic imaging methods are limited in that they cannot provide accurate information on the state of the specific tissue or tissue component undergoing interventional therapy (i.e., whether the tissue is in its native or a denatured state, necrotic or viable). Accordingly, there remains a need for a diagnostic imaging method that can non-invasively and accurately monitor the state of a specific tissue or tissue component, which can optionally provide rapid feedback of induced tissue necrosis during interventional therapies.